Tubular organs and structures such as blood vessels, the esophagus, intestines, endocrine gland ducts and the urethra are all subject to strictures i.e., a narrowing or occlusion of the lumen. Strictures can be caused by a variety of traumatic or organic disorders and symptoms can range from mild irritation and discomfort to paralysis and death. Treatment is site specific and varies with the nature and extent of the occlusion.
Life threatening stenoses are most commonly associated with the cardiovascular system and are often treated using percutaneous transluminal coronary angioplasty (PTCA). This process reduces the stricture by expanding the artery's diameter at the blockage site using a balloon catheter. However, three to six months after PTCA, approximately 30% to 40% of patients experience restenosis. Injury to the arterial wall during PTCA is believed to be the initiating event causing restenosis and primarily results from vascular smooth muscle cell proliferation and extracellular matrix secretion at the injured site. Restenosis is also a major problem in non-coronary artery disease including the carotid, femoral, iliac, popliteal and renal arteries.
Stenosis of non-vascular tubular structures is often caused by inflammation, neoplasm and benign intimal hyperplasia. In the case of esophageal and intestinal strictures, the obstruction can be surgically removed and the lumen repaired by anastomosis. The smaller transluminal spaces associated with ducts and vessels may also be repaired in this fashion; however, restenosis caused by intimal hyperplasia is common. Furthermore, dehiscence is also frequently associated with anastomosis requiring additional surgery which can result in increased tissue damage, inflammation and scar tissue development leading to restenosis.
Problems with diminished urine flow rates are common in aging males. The most frequent cause is benign prostatic hypertrophy (BPH). In this disease the internal lobes of the prostate slowly enlarge and progressively occlude the urethral lumen. A number of therapeutic options are available for treating BPH. These include watchful waiting (no treatment), several drugs, a variety of so-called “less invasive” therapies, and transurethral resection of the prostate (TURP)—long considered the gold standard.
Urethral strictures are also a significant cause of reduced urine flow rates. In general, a urethral stricture is a circumferential band of fibrous scar tissue which progressively contracts and narrows the urethral lumen. Strictures of this type may be congenital or may result from urethral trauma or disease. Strictures were traditionally treated by dilation with sounds or bougies. More recently, balloon catheters became available for dilation. Surgical urethrotomy is currently the preferred treatment, but restenosis remains a significant problem.
Recent advances in biomedical engineering have led to the development of stenting i.e., mechanical scaffolding, to prevent restenosis and keep the previously occluded lumens open. There are two general types of stents: permanent and temporary. Temporary stents can be further subdivided into removable and absorbable.
Permanent stents are used where long term structural support or restenosis prevention is required, or in cases where surgical removal of the implanted stent is impractical. Permanent stents are usually made from metals such as Phynox, 316 stainless steel, MP35N alloy, and superelastic Nitinol (nickel-titanium).
Stents are also used as temporary devices to prevent closure of a recently opened urethra following minimally invasive procedures for BPH which typically elicit post treatment edema and urethral obstruction. In these cases, the stent will typically not be covered with tissue (epithelialized) prior to removal.
Temporary absorbable stents can be made from a wide range of synthetic bio-compatible -polymers depending on the physical qualities desired. Representative bio-compatible polymers include polyanhydrides, polycaprolactone, polyglycolic acid, poly-L-lactic acid, poly-D-L-lactic acid and polyphosphate esters.
Stents are designed to be deployed and expanded in different ways. A stent can be designed to self expand upon release from its delivery system, or it may require application of a radial force through the delivery system to expand the stent to the desired diameter. Self expanding stents are typically made of metal and are woven or wound like a spring. Synthetic polymer stents of this type are also known in the art. Self-expanding stents are compressed prior to insertion into the delivery device and released by the practitioner when correctly positioned within the stricture site. After release, the stent self expands to a predetermined diameter and is held in place by the expansion force or other physical features of the device.
Stents which require mechanical expansion by the surgeon are commonly deployed by a balloon-type catheter. Once positioned within the stricture, the stent is expanded in situ to a size sufficient to fill the lumen and prevent restenosis. Various designs and other means of expansion have also been developed. One variation is described in Healy and Dorfman, U.S. Pat. No. 5,670,161. Healy and Dorfman disclose the use of a bio-compatible stent that is expanded by a thermo-mechanical process concomitant with deployment.
Approximately one-third of all patients undergoing surgery, catheterization or balloon dilation to repair bulbar urethral strictures experience restenosis. In these patients the use of urethral stents has provided satisfactory relief from symptoms. (Badlani, G. H., et al., UroLume® Endourethral Prosthesis for the Treatment of Urethral Stricture Disease: Long-term Results of the North American Multicenter UroLume® Trial. Urology: 45:5, 1993). Currently, urethral stents are composed of bio-compatible metals woven into a tubular mesh or wound into a continuous coil and are inserted endoscopically after opening the stricture by urethrotomy or sequential dilation. The stent is initially anchored in place through radial force as the stent exerts expansion pressure against the urethral wall. With woven stents epithelial cells lining the urethra begin to grow through the stent's open weave between six and 12 weeks after insertion, thereby permanently securing the stent.
For most patients this is a one time process without complication. However, some men experience post insertion complications including stent migration, excessive epithelialization, and stent encrustation. In some cases excessive epithelial tissue may be resected transurethrally. In other situations stent removal may be necessary. Depending on the condition of the stent, removal procedures range from a relatively simple transurethral procedure to open surgery with excision and urethroplasty. All complications increase patent discomfort and health care costs.
Recently, a number of bio-compatible, bioresorbable materials have been used in stent development and in situ drug delivery development. Examples include U.S. Pat. No. 5,670,161 (a thermo-mechanically expanded biodegradable stent made from a co-polymer of L-lactide and ε-caprolactone), U.S. Pat. No. 5,085,629 (a bioresorbable urethral stent comprising a terpolymer of L-lactide, glycolide and ε-caprolactone) U.S. Pat. No. 5,160,341 (a resorbable urethral stent made from polylactic acid or polyglycolic acid), and U.S. Pat. No. 5,441,515 (a bio-erodible drug delivery stent and method with a drug release layer).
The bioresorbable stents discussed in these earlier references are all designed and made from co-polymers, which is in sharp contrast to the use of the blending process of the present invention. The blending aspect of the present invention overcomes disadvantages associated with the prior art co-polymers insofar as it is more cost effective than co-polymerization, which typically must be out-sourced by end product stent manufacturers. The blending process also offers greater versatility insofar as the raw materials used in earlier co-polymeric stents were fixed in design and physical qualities. Any changes in the polymer formulation necessary to improve stent performance using a co-polymerization process can only be accomplished by having new copolymer materials manufactured by the supplier. This often results in excessive delays in product development and significantly increases research and development costs.
Furthermore, co-polymers of L-lactide and ε caprolactone are typically mostly amorphous and may be more susceptible to hydrolytic decomposition than a blend of poly-L-lactide and poly-ε-caprolactone of similar composition. Additionally, it is more difficult to maintain consistency in the manufacture of co-polymers than homopolymers, resulting in significant batch to batch variation in copolymers.
Consequently, there remains a need for a self expanding stent with stable and predictable physical characteristics suited for a wide variety of physiological conditions. In particular, there is a need for a stent making process and stent design that can be easily and cost effectively implemented for any number of application requirements.